MRI is an effective, non-invasive magnetic imaging technique for generating sharp images of the internal anatomy of the human body, which provides an efficient means for diagnosing disorders such as neurological and cardiac abnormalities and for spotting tumors and the like. Briefly, the patient is placed within the center of a large superconducting magnetic that generates a powerful static magnetic field. The static magnetic field causes protons within tissues of the body to align with an axis of the static field. A pulsed radio-frequency (RF) magnetic field is then applied causing the protons to begin to precess around the axis of the static field. Pulsed gradient magnetic fields are then applied to cause the protons within selected locations of the body to emit RF signals, which are detected by sensors of the MRI system. Based on the RF signals emitted by the protons, the MRI system then generates a precise image of the selected locations of the body, typically image slices of organs of interest.
However, MRI procedures are problematic for patients with implantable medical devices such as pacemakers and ICDs. One of the significant problems or risks is that the strong RF fields of the MRI can induce currents through the lead system of the implantable device into the tissues resulting in Joule heating in the cardiac tissues around the electrodes of leads, potentially damaging adjacent tissues. Indeed, in worst-case scenarios, the temperature at the tip of an implanted lead has been found to increase as much as 70 degrees Celsius (C) during an MRI tested in a gel phantom in a non-clinical configuration. Although such a dramatic increase is probably unlikely within a clinical system wherein leads are properly implanted, even a temperature increase of only about 80-13° C. might cause myocardial tissue damage.
Furthermore, any significant heating of cardiac tissues near lead electrodes can affect the pacing and sensing parameters associated with the tissues near the electrode, thus potentially preventing pacing pulses from being properly captured within the heart of the patient and/or preventing intrinsic electrical events from being properly sensed by the device. The latter might result, depending upon the circumstances, in therapy being improperly delivered or improperly withheld. Another significant concern is that any currents induced in the lead system can potentially generate voltages within cardiac tissue comparable in amplitude and duration to stimulation pulses and hence might trigger unwanted contractions of heart tissue. The rate of such contractions can be extremely high, posing significant clinical risks on patients. Therefore, there is a need to reduce heating in the leads of implantable medical devices, especially pacemakers and ICDs, and to also reduce the risks of improper tissue stimulation during an MRI, which is referred to herein as MRI-induced pacing.
A variety of techniques have been developed. See, for example, the following patents and patent applications: U.S. Pat. Nos. 6,871,091, 6,930,242, 6,944,489, 6,971,391, 6,985,775; U.S. Patent Application Nos. 2003/0083723, 2003/0083726, 2003/0144716, 2003/0144718, and 2003/0144719, and 2006/0085043; as well as the following PCT documents WO 03/037424, WO 03/063946, WO 03/063953. At least some of these techniques are directed to the use of RF filters, such as inductive filters or LC filters, within the leads for use in filtering RF signals induced by MRIs.
However, issues arise in determining optimal values for L and C for use within the RF filters mounted within device leads. Preferably, L and C values are selected so that the RF filter resonates at the frequency of RF signals induced by the MRI, such as at a resonant frequency of, e.g., 65.3 MHz of 1.5 T or 128 MHz of 3 T. In principle, a wide range of L and C values can potentially be used to achieve a desired resonant frequency based on: 2π*f=1/sqrt(L*C) wherein “f” is the resonant frequency. (Solving for L, this equation may also be represented as: L=1/(4π2f2C).) However, it has been discovered that not all combinations of permissible L and C values are equally effective in reducing lead heating.
FIG. 1 illustrates a range of L and C values for an LC filter wherein curve 1 represents combinations of values satisfying the aforementioned equations for a resonant frequency of 65.3 MHz. Any combination of L and C values along the curve satisfies the equations, such as combination LC1 (where L=10 nH and C=620 pF) or combination LC2 (where L=297 nH and C=20 pF). However, lead modeling and in vitro tests have shown that combination LC2 achieves a significantly greater reduction in tip heating as compared to combination LC1.
FIG. 2 illustrates expected increases in lead temperatures arising within a lead due to an 65.3 MHz/1.5 Tesla MRI for one in vitro lead implementation and three non in vitro lead implementations, with the expected temperature increases determined via computer modeling. In FIG. 2, the expected temperature increase within the lead is shown for various exemplary lead lengths in the range of 30-45 cm. Curve 2 represents the expected increase in lead tip temperature for a non in vitro lead during the MRI while using an LC filter with parameters LC2. Curve 3 represents the expected increase in lead tip temperature for a non in vitro lead during the MRI while using an LC filter with parameters LC1. Curve 4 represents the expected increase in lead tip temperature for a non in vitro lead during the MRI without an LC filter. Curve 5 represents the expected increase in lead tip temperature for an in vitro lead during an MRI, also without an LC filter. As can be seen, without an LC filter, the tip temperature is expected to increase by as much as 20 degrees C. due to the MRI (with the greatest increase expected in a 30 cm lead). Although LC filter combination LC1 is expected to achieve a reduction in tip temperature during the MRI as compared to the control lead, the LC2 filter combination is expected to achieve a much greater reduction in tip temperatures. Indeed, with the LC2 combination, only a minimal temperature increase is expected during an MRI, regardless of lead length.
FIG. 3 illustrates increases in lead temperature measured within in vitro test leads exposed to an 65.3 MHz/1.5 T MRI. Again, temperature increases during the MRI are plotted for the two exemplary LC combinations (LC1 and LC2) and for a control lead without an LC filter. In FIG. 3, the measured temperature increase within the lead is shown for various exemplary lead lengths in the range of 25-53 cm. Curve 6 represents the increase in lead tip temperature measured during the MRI when employing an LC filter with parameters LC2. Curve 7 represents the increase in lead tip temperature measured when employing an LC filter with parameters LC1. Curve 9 represents the increase measured in lead tip temperature without an LC filter. As can be seen, without an LC filter, the tip increased by as much as 33 degrees C. due to the MRI (with the greatest increase occurring in a 25 cm lead). LC filter combination LC1 achieved only minimal reductions in tip temperature during the MRI as compared to the control lead. However, the LC2 filter combination achieved significant reductions in tip temperatures, up to 90% or more. Indeed, the in vitro results of FIG. 3 demonstrate that an even more significant reduction in tip temperature can be achieved when using the LC2 combination than the LC1 combination, as compared to the reduction shown in the modeling results of FIG. 2 for non in vitro lead models.
As can be appreciated, there is a need to understand the significant differences in lead temperature reduction achieved with different combinations of resonant LC values in practical implementations. Furthermore, there is a need to determine suitable combinations of L and C values for LC filters for use within particular implementations so as to achieve an effective reduction in lead temperature during an MRI, as compared to leads without LC filters. It is to these ends that the invention is generally directed.